Split energy level radiation detection

ABSTRACT

An energy discriminating apparatus and method is disclosed for use in connection with digital radiography and fluoroscopy. In use of the detection system and method an x-ray source is actuated to direct x-rays through a patient&#39;s body, the x-rays including both higher and lower energy radiation. A first detector element, including a plurality of segments, is positioned opposite the source to receive and respond predominantly to x-rays in a lower energy range, the remaining x-rays, being generally of higher energy, passing through the first detector element. A second detector element, also including a plurality of segments, each segment including a phosphor coating layer and a sensor, is positioned to receive and respond to the higher energy radiation passing through the first element. The sensors are coupled respectively to each detector element segment for substantially simultaneously sensing the response and spatial location, relative to the detector elements, of radiation to which each detector element respectively responds. A filter element is interposed between the first and second detectors to enhance discrimination in the energy response of the respective detector elements. Particular preferred detector phosphor materials are identified. The sensors produce separately and simultaneously information representing patterns of relatively lower and higher energy emergent from the patient&#39;s body. Digital data processing and conversion equipment responds to the sensors to produce digital information representing each of said images, which can be digitally processed to enhance image characteristics.

DESCRIPTION

1. Technical Field

This invention relates to the field of medical diagnostic imaging andmore particularly to an improved x-ray detector for use in digitalradiography and fluoroscopy. The detector provides separate simultaneousrepresentations of different energy radiation emergent from a subject.

2. Background Art

Radiography and fluoroscopy are long well known diagnostic imagingtechniques.

In a conventional radiography system, an x-ray source is actuated todirect a divergent area beam of x-rays through a patient. A cassettecontaining an x-ray sensitive phosphor screen and film is positioned inthe x-ray path on the side of the patient opposite the source. Radiationpassing through the patient's body is attenuated in varying degrees inaccordance with the various types of tissue through which the x-rayspass. The attenuated x-rays from the patient emerge in a pattern, andstrike the phosphor screen, which in turn exposes the film. The x-rayfilm is processed to yield a visible image which can be interpreted by aradiologist as defining internal body structure and/or condition of thepatient.

In conventional fluoroscopy, a continuous or rapidly pulsed area beam ofx-rays is directed through the patient's body. An image intensifier tubeis positioned in the path of the beam opposite the source with respectto the patient. The image intensifier tube receives the emergentradiation pattern from the patient, and converts it to a small,brightened visible image at an output face. Either a mirror or closedcircuit television system views the output face and produces a dynamicreal time visual image, such as on a CRT, a visual image forinterpretation by a radiologist.

More recently, digital radiography and fluoroscopy techniques have beendeveloped. In digital radiography, the source directs x-radiationthrough a patient's body to a detector in the beam path beyond thepatient. The detector, by use of appropriate sensor means, responds toincident radiation to produce analog signals representing the sensedradiation image, which signals are converted to digital information andfed to a digital data processing unit. The data processing unit records,and/or processes and enhances the digital data. A display unit respondsto the appropriate digital data representing the image to convert thedigital information back into analog form and produce a visual displayof the patient's internal body structure derived from the acquired imagepattern of radiation emergent from the patient's body. The displaysystem can be coupled directly to the digital data processing unit forsubstantially real time imaging, or can be fed stored digital data fromdigital storage means such as tapes or discs representing patient imagesfrom earlier studies.

Digital radiography includes radiographic techniques in which a thin fanbeam of x-ray is used, and other techniques in which a more widelydispersed so-called "area beam" is used. In the former technique, oftencalled "scan (or slit) projection radiography" (SPR) a fan beam of x-rayis directed through a patient's body. The fan is scanned across to thepatient, or the patient is movably interposed between the fan beam x-raysource and an array of individual cellular detector segments which arealigned along an arcuate or linear path. Relative movement is effectedbetween the source-detector arrangement and the patient's body, keepingthe detector aligned with the beam, such that a large area of thepatient's body is scanned by the fan beam of x-rays. Each of thedetector segments produces analog signals indicating characteristics ofthe received x-rays.

These analog signals are digitized and fed to a data processing unitwhich operates on the data in a predetermined fashion to actuate displayapparatus to produce a display image representing the internal structureand/or condition of the patient's body.

In use of the "area" beam, a divergent beam of x-ray is directed throughthe patient's body toward the input face of an image intensifier tubepositioned opposite the patient with respect to the source. The tubeoutput face is viewed by a television camera. The camera video signal isdigitized, fed to a data processing unit, and subsequently converted toa tangible representation of the patient's internal body structure orcondition.

One of the advantages of digital radiography and fluoroscopy is that thedigital image information generated from the emergent radiation patternincident on the detector can be processed, more easily than analog data,in various ways to enhance certain aspects of the image, to make theimage more readily intelligible and to display a wider range ofanatomical attenuation differences.

An important technique for enhancing a digitally represented image iscalled "subtraction". There are two types of subtraction techniques, onebeing "temporal" substraction, the other "energy" subtraction.

Temporal, sometimes called "mask mode" subtraction, is a technique thatcan be used to remove overlying and underlying structures from an imagewhen the object of interest is enhanced by a radiopaque contrast agent,administered intra-arterially or intra-venously. Images are acquiredwith and without the contrast agent present and the data representingthe former image is subtracted from the data representing the latter,substantially cancelling out all but the blood vessels or anatomicalregions containing the contrast agent. Temporal subtraction is,theoretically, the optimum way to image the enhancement caused by anadministered contrast agent. It "pulls" the affected regions out of aninterfering background.

A principle limitation of digital temporal subtraction is thesusceptibility to misregistration, or "motion" artifacts caused bypatient movement between the acquisition of the images with and withoutthe contrast agent.

Another disadvantage of temporal subtraction is that it requires the useof a contrast material and changes in the contrast caused by the agentmust occur rapidly, to minimize the occurrence of motion causedartifacts by reducing the time between the first and second exposureacquisition. Temporal subtraction is also not useful in studiesinvolving rapidly moving organs such as the heart. Also, theadministration of contrast agents is contraindicated in some patients.

An alternative to temporal subtraction, which is less susceptible tomotion artifacts, is energy subtraction Whereas temporal subtractiondepends on changes in the contrast distribution with time, energysubtraction exploits energy-related differences in attenuationproperties of various types of tissue, such as soft tissue and bone.

It is known that different tissues, such as soft tissue (which is mostlywater) and bone, exhibit different characteristics in their capabilitiesto attenuate x-radiation of differing energy levels.

It is also known that the capability of soft tissue to attenuatex-radiation is less dependent on the x-ray's energy level than is thecapability of bone to attenuate x-rays. Soft tissue shows less change inattenuation capability with respect to energy than does bone.

This phenomenon enables performance of energy subtraction. In practicingthat technique, pulses of x-rays having alternating higher and lowerenergy levels are directed through the patient's body. When a lowerenergy pulse is so generated, the detector and associated digitalprocessing unit cooperate to acquire and store a set of digital datarepresenting the image produced in response to the lower energy pulse. Avery short time later, when the higher energy pulse is produced, thedetector and digital processing unit again similarly cooperate toacquire and store a set of digital information representing the imageproduced by the higher energy pulse. The values obtained representingthe lower energy image are then subtracted from the values representingthe higher energy image.

Since the attenuation of the lower energy x-rays by the soft tissue inthe body is approximately the same as soft tissue attenuation of thehigher energy x-rays, subtraction of the lower energy image data fromthe higher energy image data approximately cancels out the informationdescribing the configuration of the soft tissue. When this informationhas been so cancelled, substantially all that remains in the image isthe representation of bone. In this manner, the contrast and visibilityof the bone is substantially enhanced by energy subtraction.

Energy subtraction has the advantage, relative to temporal subtraction,of being substantially not subject to motion artifacts resulting fromthe patient's movement between exposures. The time separating the lowerand higher energy image acquisitions is quite short, often less than onesixtieth of a second.

Details of energy subtraction techniques in digital radiography andfluoroscopy are set forth in the following technical publications, allof which are hereby incorporated specifically by reference:

Hall, A.L. et al: "Experimental System for Dual Energy ScannedProjection Radiology". Digital Radiography proc. of the SPIE 314:155-159, 1981;

Summer, F.G. et al: "Abdominal Dual Energy Imaging" Digital Radiographyproc. SPIE 314: 172-174, 1981;

Blank, N. et al: "Dual Energy Radiography: a Preliminary Study". DigitalRadioqraphy proc. SPIE 314: 181-182, 1981; and

Lehman, L.A. et al: "Generalized Image Combinations in Dual kVp DigitalRadiography", Medical Physics 8: 659-667, 1981.

Dual energy subtraction has been accomplished, as noted above, bypulsing an x-ray source in a digital scanning slit device at two kVp's,typically 120 and 80 kVp, and sychronizing the pulses with a rotatingfilter which hardens the high kVp pulses by filtering out the lowerenergy x-ray. This results in the patient and x-ray detectorsequentially seeing high energy and low energy beams from which the massper unit area of bone and soft tissue can be solved for.

In energy subtraction, it is desirable that the two energy levels shouldbe widely separated. This is necessary in order to accurately define themasses per unit area of bone and soft tissue.

With a slit scanning device, such as described above, sequentiallypulsing the x-ray tube at 120 and 80 kVp is technically difficult andgives rise to very difficult problems in a practical clinical device.The switching frequency has to be on the order of 500 Hz. andinsufficient photons (x-ray energy per pulse) results when the highestcapacity x-ray tubes are combined with realistically narrow slit widthsand scanning times.

In connection with CT (computerized tomography) applications, a twolayer energy sensitive detector has been proposed. In this proposal, afirst calcuim fluoride layer is provided for sensing lower level x-rayradiation, and a second downstream sodium iodide layer senses higherenergy radiation passing through the first layer. Light caused byradiation in each of the two layers is separately sensed by respectivephotomultiplier tubes.

DISCLOSURE OF THE INVENTION

The disadvantages and problems of the prior art are alleviated oreliminated by the use of an energy discriminating radiation detectorincluding three elements. The detector includes a first elementpredominantly responsive to radiation of a first energy range, and asecond element positioned behind the first, responsive to radiation in asecond and higher energy range, along with a radiation filter interposedbetween the first and second elements.

Thus, an energy sensitive x-ray detector system for use in digitalradiography is provided. For each picture element of the radiographicprojection, the detector provides two readings from which the mass perunit area of bone and soft tissue through which the x-ray beam passescan be determined.

The energy sensitive x-ray detector employs a low atomic number phosphorscreen or discrete array of phosphor segments coupled to a photodiodearray, followed by a high atomic number phosphor screen or discretesegment array similarly coupled.

An energy sensitive segment of an element of the detector systemconsists of a low atomic number phosphor coating layer coupled to afirst photodiode, followed by a high atomic number phosphor coatinglayer coupled to a second photodiode. The low atomic number phosphorpreferentially absorbs the low energy photons emerging from the patientand transmits most of the higher energy photons, a large percentage ofwhich are absorbed in the second (higher atomic number) phosphor.

Placing an appropriate filter between the two phosphor/photodiode arraysincreases or hardens the effective energy of the x-ray spectrum incidenton the second phosphor and results in a greater and more desirableenergy separation between the x-ray spectra absorbed in the two phosphorlayers.

In accordance with another embodiment, a split energy radiation detectoris provided including a first energy responsive element comprising aquantity of phosphor material including one of yttrium oxysulfide andzinc cadmium sulfide, and a second energy responsive element positionedto receive energy passing through said first element, said secondelement including one of gadolinium oxysulfide and cadmium tungstate.

In accordance with another specific aspect of the invention, theradiation filter interposed between the two elements or layers is madeof a material containing copper.

In accordance with a broader aspect of the invention, there is provideda split energy radiation detector screen comprising a deck of separatedetector elements at least partially mutually superposed, each elementbeing capable of producing information spatially locating radiationincident on the screen.

These and other aspects of the present invention will become moreapparent from a consideration of the following description and of thedrawings, in which:

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a plan pictorial and block illustration of a systemincorporating the present invention;

FIGS. 1A-1E are detail views illustrating a portion of the system ofFIG. 1;

FIG. 2 is a side view illustrating a portion of the system illustratedin FIG. 1;

FIG. 2A is a detailed side view illustrating a portion of the system ofFIG. 1;

FIG. 3 is a perspective view of an alternate embodiment of a portion ofthe system of FIG. 1.

FIG. 3A is a graphical description of a preferred feature of the portionof the system illustrated in FIG. 2.

FIG. 4 is a graphical representation of operating characteristics of theportion of the system illustrated in FIG. 2;

FIG. 5 is a block diagram illustrating another system incorporating anembodiment of the present invention.

BEST MODE FOR CARRYING OUT THE INVENTION

FIG. 1 illustrates a slit projection type of digital radiography systemin which the present invention is incorporated. The system S scans athin fan beam of multi-energetic x-rays over a patient's chest andseparately detects a pattern of x-rays of different energies emergentfrom the patient's body. Information represented by the detected x-raysis processed and displayed to illustrate a representation of an image ofthe patient's internal body structure or condition.

More specifically, the system S includes an x-ray source X affixed tomounting structure M for projecting a thin fan beam B of x-rays throughthe body of a patient P, to strike an aligned array D of detectorsegments. The fan beam B is confined by a forward slit K tosubstantially a vertical plane. The detector array D constitutes avertical stack of individual detector segments E, described in moredetail below, and aligned with the vertical plane defined by the beam B.An aft slit J attached to the detector D receives and aids in thedefinition of the planar beam B.

The x-ray source X is mounted on the structure M to rotate about avertical axis, defined in FIG. 1 as extending into the paper. Mechanicallinkage L couples the x-ray tube X to the detector array D and causesthe detector array D to scan behind the patient's body in the directionsof the arrows A, A¹, in order to maintain the detector D aligned withthe beam B throughout the scanning rotative motion of the x-ray tube X.

The x-ray source X is controlled to emit either a continuous beam or arapid succession of x-ray pulses in the form of the fan beam B. Thex-ray tube X and the detector D are synchronously scanned, about avertical axis, across the patient from one side of his body to theother. The detector output is periodically sampled. Each samplingproduces signals representing a line of image information. Over thecourse of the scan from side to side, signals are developed describing aplurality of lines, which together constitute an area image of thepatient's internal body structure.

Details of some aspects of a digital radiography system such asdescribed above are set forth in the following publications, herebyexpressly incorporated by reference.

Arnold, B.A. et al, "Digital Radiography: An Overview" Proc. of S.P.I.E.Vol. 273, Mar. 1981;

Kruger, R.A. et al, "A Digital Video Image Processor for Real Time X-RaySubtraction Imaging" Optical Engineering Vol. 17 No. 6 (1978).

The detector D separately detects x-rays of different energy rangesimpinging on the detector array. An element of the detector array, byway of two sets of leads 01, 02, transmits analog signals representingdetected x-rays within lower and higher energy ranges, respectively.

The signals on the lead sets 01, 02, are provided to ananalog-to-digital converter C which digitizes the outputs and feeds themto a digital processing and receiving unit DPU. The DPU processes thesedigitized output signals to construct a digital representation of animage of the patient's internal body structure scanned by the x-ray beamB, on a line-by-line basis. Digital signals from the DPU are convertedto analog form by way of a digital-to-analog converter, and fed to adisplay unit T, which in response, produces an image in visual formcorresponding to the image representing signals from the DPU.

Optionally, digital storage means can be provided in conjunction withthe DPU in order to digitally store the image representations for futureuse. In such event, the digitally stored signals can be played blackthrough the DPU, converted to analog form, and their correspondingimages displayed at a later time on the display apparatus T.

FIGS. 1A and 1B illustrate (in simplified form, for clarity) particularconfigurations of the face of the detector array D, as viewed from theright in FIG. 1. In FIG. 1A, for example, it is seen that the detector Dcomprises a linear vertically stacked elongated array of detectorsegments E.

An alternative embodiment to the vertical linear detector array shown inFIG. 1A is illustrated in FIG. 1B. This is known as a "staggered" array.The staggered array consists of two side-by-side vertical columns ofdetector segments E, E¹. One of the vertical columns, however, isslightly vertically displaced with respect to the other, by a distanceequal to one-half the height of a single detector segment.

FIGS. 1C-1E illustrate in simplified form several embodiments of thedetector configuration of FIG. 1A as viewed from the right side in FIG.1A. FIGS. 1C-1E, however, are not intended to show the dual layeredstructure of the detector segments, which will be later discussed indetail, such as in connection with FIG. 2A. The detector arrays aredivided into individual segments in one of three ways. In oneembodiment, shown in FIG. 1C, the detector array D comprises anelongated vertical screen strip 10 of particles of radiation sensitivematerial which are glued together with a binder and affixed to a backingof a suitable material such as polyester. The radiation sensitivematerial respnds to incident radiation to produce light. Behind theradiation sensitive screen 10 is a vertical array of adjacentphotodiodes 12. Each photodiode responds to radiation-caused light inthe screen 10 to produce an analog electrical signal indicatingbrightness of the flash caused by the sensed radiation events. Each ofthe photodiodes 12 responds primarily to light from radiation eventsoccurring within a portion of the screen material 10 located adjacentthe photodiode.

Special "cellularized" detector configurations are illustrated in FIGS.1D and 1E. Cellularized detectors have the advantage of reducing theeffects of energy scatter within the detector array.

In the form illustrated in FIG. 1D, the detector screen 10 is grooved asillustrated for example at reference character 14, and the grooves areimpregnated with a reflective material, such as aluminum oxide, toreduce the effects of light within the screen 10. The grooves arealigned with the junctions between each of the adjacent photodiodes 12.

Another form of cellularized detector arrangement is illustrated in FIG.1E. In that embodiment, rather than utilizing an homogeneous screen,with or without grooves, separate crystalline portions 16 of radiationsensitive material are employed. Each crystal is matched to an adjoiningphotodiode and separated from adjacent crystals by a reflective layer.The size of each of the crystals corresponds to the size of itsadjoining photodiode 12.

In all of the foregoing detector arrangements, the photodiodes areadhered to the screen portion 10 by a mechanical pressing operation,which can optionally be aided by a small quantity of adhesive, and/or asmall amount of optical coupling grease to enhance the degree of opticalcoupling between the screen 10, or crystals 16, and the photodioes 12.

As pointed out above, it is desirable, when practicing the energysubtraction image processing technique, to be able to separatelyrepresent different energy radiation which impinges on the detectorsegments. Herein is disclosed a particular dual layered, energydiscriminating structure for each detector segment which facilitatesachievement of this goal.

FIG. 2 illustrates a particular layered detector segment structure foruse as a component of an energy sensitive radiation detector array D.The detector responds to radiation incident upon it, transmitted in adownward direction with respect to FIG. 2, to produce two outputs atleads 18, 20. The output at lead 18 represents radiation incident uponthe detector segment having an energy level in a lower energy range. Theoutput at the lead 20 represents the detector segment's response toincident x-ray radiation having an energy level in a second, higherenergy range.

The detector segment includes a first elemental layer 22 primarilyresponsive to lower energy x-rays, and a second elemental layer 24responsive to higher energy x-rays. Each of the layers 22, 24, includesa phosphor coating layer 26, 28, respectively, and a photodiode 30, 32,each respectively optically coupled to the phosphor layers 26, 28.

The choice of materials for the phosphor layers 26, 28, is significant.For example, preferred phosphor material for the first phosphor layer 26include yttrium oxysulfide, and zinc cadmium sulfide. Alternativephosphors are barium sulfate, barium cadmium sulfate, lanthimumoxysulfide and barium fluorochloride.

For the second phosphor layer 28, preferred phosphors are gadoliniumoxysulfide and cadmium tungstate. Alternative phosphor materials for thephosphor layer 28 include calcium tungstate and barium lead sulfate.

A preferred phosphor coating weight for the first phosphor layer 26 isabout 20 to 100 milligrams (mg) per square centimeter (cm²).

Preferred phosphor coating weights for the second phosphor layer lie inthe range from approximately 50 to 1000 mg/cm².

The detector segment described above as embodying this invention isuseful not only in linear detector element arrays such as used in scanor slit projection radiography, but also in larger area detector screensused in digital radiography systems incorporated divergent, "area" x-raybeams. In the latter case, a phosphor matrix embodying the detector canconsist of either a single integral x-ray intensifying screen, acellularized intensifying screen, or a cellularized matrix of individualphosphor crystals.

The segments have equal square dimensions in each layer.

The dimensions of the individual cell segments, where a cellularizedstructure is used, are equal to the photodiode matrix array spacing,such that each individual photodiode is congruent with its cell segment.

The cell segment dimensions are greater in the second layer of thedetector than in the first. The relationship between cell segmentdimensions in the first and second layers is expressed by the following:

    (D2/D1)=(F2/F1)

where

D2=the second detector photodiode, dimension;

D1=the first detector photodiode dimension;

F2=the distance from the x-ray source focal spot to the second detectorlayer 24, and

F1=the distance from the x-ray focal spot to the first detector layer 22(see FIG. 3A for a graphical illustration of these values).

This relation applies irrespective of whether a slit projection or areascreen is employed.

It is desirable that the phosphor material selected for the firstphosphor layer 26 have a primary absorber atomic number lying in therange of 39 to 57. The corresponding desirable atomic number range forthe phosphor materials' primary absorber selected for the second layer28 is 56 to 83.

The capability of the detector structure of this invention todistinguish between incident x-rays of differing energy ranges can beenhanced by the interposition of a filter layer 36 between the first andsecond layers 22, 24. A preferred filter material is one containingcopper, such as brass. A preferred filter thickness, where brass isused, is approximately 0.5 millimeters (mm). The range of practicalbrass filter thicknesses is from about 0.2 mm to about 1.0 mm.Alternative filters can comprise either single or multiple filterelements made of material ranging in atomic number from approximately 24to 58.

When a detector element constructed in accordance with the presentlyindicated preferred embodiment is used, a desirable energy spectrum forthe x-ray source is from about 80 kVp to 150 kVp, or even higher, iftube technology permits.

The degree of spacing between the first and second layers 22, 24 of thedetector segment is not particularly critical. Spacing between the firstand second layers can suitably vary from almost physical contact toabout 3 or more centimeters (cm). The spacing between the filter layer36 and the first and second layers 22, 24 is not critical either.

As mentioned above, figures such as FIG. 1C show a side view of thedetector array D in a form simplified for clarity. FIG. 1C is simplifiedin that it shows only one of the two detector elements or layers whicheach contain a plurality of detector segments as defined by thedimensions of the photodiodes 12.

FIG. 2A is provided to show the dual detector element (layer) structurewhich is the present subject. FIG. 2A shows how the detailed structureof FIG. 2 appears, when incorporated into a linear detector array D.FIG. 2A represents a side view of such an array.

FIG. 2A illustrates the two detector elements or layers 22, 24 onepositioned behind the other with respect to the incident radiation fromthe source. Each element includes respectively a coating layer ofphosphor 26, 28, and a set of photodiodes respectively indicated at 30,32. Between the elements is located the filter element 36.

Each photodiode has a lead emergent therefrom for transmitting itsanalog radiation indicating signal to the appropriate one of the leadgroups 01, 02, as described generally above. For purposes of clarity,only representative leads are shown in FIG. 2A.

The application of the split energy radiation detector of this inventionis by no means limited to a linear array of detectors, for use in slitprojection digital radiography, the environment described in detailabove. The present invention can also be embodied in a so-called "area"detector, i.e., a relatively large rectangular radiation detectorcovering a relatively expansive portion of the patient's body, designedfor use with so-called "area" beams, which diverge from the source toexpose the radiation detector simultaneously over its entire face. Onelayer of such an area detector is illustrated in FIG. 3, it beingunderstood that such an area detector includes two such layers, onebehind the other.

Other types of area detectors exist in which use of this invention isadvantageous. One such area detector includes a first phosphor layer ofrelatively low atomic number, as described above, coupled to aradiographic film layer, behind which is a second higher atomic numberphosphor screen coupled to a second piece of film. Also, instead of thefilm portions, photoconductive or thermoluminescent plates could beused.

The principles analogous to the construction of the cellularized anduncellularized detectors described above in conjunction with FIGS. lAthrough lE can also be applied to area detectors as well.

Where such an area detector is used, the decoding electronics forlocating the sites of radiation events across the face of an areadetector are more complicated than in the case of the linear detectorarray discussed above. Details of a system for accomplishing this, whichcould analogously be applied to an area detector embodying thisinvention, are set forth in a publication entitled "A Practical GammaRay Camera System Using High Purity Germanium" published in the February1974 issue of IEEE Trans Nuc Sci and prepared by the Ohio StateUniversity Department of Nuclear Engineering under the auspices of aNational Institute of Health contract. This publication is expresslyincorporated by reference herein.

As may be implied by the above incorporated publication, the presentinvention is applicable to radiation detector technology employing otherthan phosphor materials which convert radiation events into lightenergy. The principles of this invention can be incorporated as wellinto radiation detection technology utilizing other types of radiationsensitive material, such as solid state materials which convert incidentradiation into electrical signals which represent radiation incident onthe material, without the need for converting such energy to the form oflight.

Energy Sensitive Experiment and Results

The arrangement of the first and second detector layers employed in theexperiment was in effect as shown in FIG. 2. A Lucite and aluminumphantom 38 was employed to simulate soft tissue and bone. Theexperimental results are tabulated in Table 1 for a typical 120 kVpradiation level and plotted in FIG. 4. Note how the iso-Lucite andiso-aluminum lines are more distinct when the brass filter is insertedbetween the first and second detector layers. From the data in Table 1the relative uncertainty in estimating the thickness of Lucite andaluminum can be calculated and these results are tabulated in Table 2.Note that the ability to discriminate Lucite and aluminum is improvedwhen the brass filter is inserted between the first and second detector.

The first phosphor layer was a 43 mg/cm² coating of yttrium oxysulfide.The second phosphor layer was a 110 mg/cm² coating of gadoliniumoxysulfide.

                  TABLE 1                                                         ______________________________________                                        Experimental Results for a Constant,                                          Typical Exposure Level                                                        Brass  Lucite   Aluminum                                                      (cm)   (cm)     (cm)       (R.sub.1)                                                                              (R.sub.2)                                 ______________________________________                                        0      0        0          3167     3809                                              2.54    0          1662     2466                                              5.08    0           917     1451                                              8.89    0           398      679                                      10.16       0          309    2.59  529  3.18                                 11.43       0           235        415                                        10.16       .1          275        491                                        10.16       .2          249        455                                        10.16       .4         209    2.22  400  2.40                                        10.16    .8          150      308                                      .0558  0        0          3196     2293                                              2.54    0          1697     1390                                              5.08    0           945      338                                              8.89    0           408      400                                      10.16       0          312    2.47  316  3.06                                 11.43       0           242        249                                        10.16       .1          282        298                                        10.16       .2          255        278                                        10.16       .4         211    2.21  248  2.29                                 10.16       .8          154        197                                        ______________________________________                                    

                  TABLE 2                                                         ______________________________________                                        Lucite and Aluminum Discrimination for                                        10.2 cm of Lucite and 4 mm of aluminum                                        Brass filter                                                                          Lucite    Aluminum  % Lucite                                                                              % Aluminum                                Thickness                                                                             Resolution                                                                              Resolution                                                                              Resolution                                                                            Resolution                                ______________________________________                                          0 mm  0.24 cm   0.102 cm  2.4     19.0                                      0.56 mm 0.16 cm    0.05 cm  1.6     12.5                                      ______________________________________                                    

A split energy level radiation detector such as illustrated in detail inFIG. 2 is also applicable in conventional radiography systems as aphototimer. Figure 5 illustrates such a system. An x-ray source 50directs a beam 51 of x-ray through the body of a patient P and onto aconventional radiation screen 52. A split level radiation detector 54,constructed in accordance with the structure detailed in FIG. 2 ispositioned as a phototimer behind the screen to receive that portion ofthe x-ray energy from the beam 51 which passes through the screen 52.

The phototimer 54 produces, on leads 53, 55, signals indicating theamount of received energy in separate lower and higher energy ranges,respectively. These separate energy indicating signals are fed to a duallevel energy integrator 56.

The energy integrator 56 includes circuitry for separately integratingthe amount of energy, over time, indicated by the outputs on the leads53, 55.

When the integrated energy values developed by the integrator 56accumulate to a predetermined criteria, the integrator 56 produces asignal to a tube control circuit 58 which terminates operation of thesource 50 in response to the accumulation of the particularpredetermined integrated energy criterian.

The energy criterian governing the time of x-ray exposure can beselected in accordance with known principles by those with skill in theart. This criterion can be defined as the accumulation of apredetermined amount of energy in either of the sensed energy ranges, orcan be a function of both sensed energy levels.

It is to be understood that this description of one embodiment of thepresent invention is intended as illustrative, and not exhaustive, ofthe invention. It is to be further understood that those of ordinaryskill in the relevant art may make certain additions, deletions andmodifications to this embodiment of the invention as described herein,without departing from the spirit or the scope of the invention, asdescribed in the appended claims.

I claim:
 1. An energy discriminating radiation detector comprising:(a) afirst element comprising a first material of a kind which ispreferentially responsive to penetrative radiation of a first energyrange; (b) a second element comprising a second material different inkind from said first material and of a kind which is preferentiallyresponsive to penetrative radiation of a second energy range extendinghigher than said first energy range and which is positioned to receiveradiation which has penetrated through a portion of said first element;and (c) a filter of penetrative radiation interposed between said firstand second elements.
 2. The detector of claim 1, wherein said filtercontains copper.
 3. The detector of claim 1, wherein said filtercomprises brass.
 4. The detector of claim 2 or 3, wherein said filter isselected to have a thickness of from about 0.2 mm to about 1.0 mm. 5.The detector of claim 1, wherein each said element comprises:(a) aphosphor layer, and (b) a photodiode optically coupled to the phosphor.6. A split energy radiation detector comprising:(a) a first energyresponsive element comprising a layer of phosphor material including oneof yttrium oxysulfide and zinc cadmium sulfide; and (b) a second energyresponsive element positioned to receive energy penetrating through saidfirst element, said second element including a second phosphor layercomprising one of gadolinium oxysulfide and cadmium tungstate.
 7. Thedetector of claim 6 further comprising:a copper containing filterelement interposed between said first and second elements.
 8. Thedetector of claim 6, wherein:(a) said first phosphor layer has a coatingweight of about 20 to 100 mg/cm², and (b) said second phosphor layer hasa coating weight of about 50 mg/cm² to 1000 mg/cm²
 9. A digitalradiography system comprising:(a) an x-ray source for directing x-raysalong a path; (b) a split energy radiation detector spaced from thesource to receive x-rays from said source, said detector comprising:(i)a first element comprising a first material of a kind which ispreferentially responsive to radiation of a first energy range and beinglocated in said path; (ii) a first sensor for sensing radiation responseof said first element; (iii) a second element at least partiallypositioned to receive source radiation passing through said firstelement, said second element comprising a second material of a kindwhich is preferentially responsive to radiation of a second energy levelextending higher than said first range; (iv) a second sensor for sensingradiation response of said second element; and (c) interpretivecircuitry coupled to said sensors for at least partially digitizinginformation from said sensors and producing from said digitizedinformation a representation of at least a portion of internal bodystructure of a subject when interposed in said path.
 10. The system ofclaim 9, wherein said first material includes one of yttrium oxysulfideand zinc cadmium sulfide.
 11. The system of claim 9, wherein said secondmaterial includes one of gadolinium oxysulfide and calcium tungstate.12. The system of claim 9, further comprising: an x-ray filter layerbetween said first and second elements.
 13. The system of claim 12,wherein said filter layer contains copper.
 14. The system of claim 9,wherein said sensors each comprise a photodiode.
 15. The system of claim9, wherein said x-ray source is capable of simultaneously producingx-rays in both said energy ranges.
 16. The system of claim 9, whereineach of said elements is substantially planar, one said element beingsubstantially behind the other with respect to the source.
 17. Animaging method comprising the steps of:(a) directing x-rays through asubject to be imaged, said x-rays including both higher and lower energyradiation; (b) separately detecting higher and lower energy x-radiationemergent from the subject by passing said radiation successively throughscintillators comprising respectively different kinds of materials eachpreferentially responsive to radiation of a different one of said lowerand higher energy ranges, including sensing responses of saidscintillators; (c) at least partially digitizing information derived insaid detecting step; (d) processing said digitized information; and (e)utilizing said processed digital information to produce a representationof internal structure of the subject.
 18. The method of claim 17,wherein said digital processing step includes a step of subtractinginformation obtained in said lower energy sensing step from infromationobtained in said higher energy sensing step.
 19. The method of claim 17,wherein said sensing step comprises producing information in response toradiation incident on a plurality of separate detector elements, saidinformation including spatial location representation of said incidentradiation with respect to a said sensing element.
 20. An energydiscriminating radiation detecting method utilizing first and seconddetector elements, a first of said elements being preferentiallyresponsive to radiation of a first energy range, a second of saidelements being preferentially responsive to energy of a second rangeextending higher than said first energy range, said method comprisingthe steps of;(a) directing radiation extending over both said first andsecond energy ranges through a subject; (b) positioning said firstelement to receive incident radiation emergent from the subject forresponse thereto; (c) positioning said second element to receiveradiation from the source passing through said first element, and (d)filtering radiation transmitted through said first element prior to thearrival of said energy incident upon said second element.
 21. Aradiographic system comprising:(a) an x-ray source; (b) a radiationdetector positioned to receive x-rays from the source; (c) a phototimercomprising:(i) an energy discriminating detector located to receivex-rays from the source and to produce signals indicating x-ray energyreceived in each of two energy ranges, and (ii) circuitry coupledbetween the discriminating detector and the source for controlling thesource as a function of the x-rays detected in said two energy ranges.22. A radiation imaging system comprising:(a) a source of penetrativeradiation; (b) a dual energy detector assembly comprising twoside-by-side columns of individual detector elements, one column beingstaggered with respect to the other by a distance equal to less than thedimension of a single detector element taken along the direction of itscolumn, and additional detector elements positioned behind said columns,relative to said source; (c) mounting structure for maintaining saidsource and said detector assembly sufficiently spaced to provide asubject examining space and for maintaining said detector alignedcontinuously in said penetrative radiation when produced by said source;(d) power means for actuating said source to direct penetrative radationthrough the subject examination space and incident onto the detectorassembly; (e) means coupled to said detector elements for producing animage of a portion of a subject, when located in the subject space, fromradiation emergent from said subject.
 23. The system of claim 22,wherein said staggered columns of detector elements are offset withrespect to one another by a distance equal approximately one-half theheight of a single detector element taken in a direction along itscolumn.
 24. An energy discriminating radiation detector comprising:(a) afirst component comprising a first material of a first kind which ispreferentially responsive to penetrative radiation of a first energyrange; (b) a second component comprising a second material different inkind from said first material and of a kind which is preferentiallyresponsive to penetrative radiation of a second energy range extendinghigher than said first energy range, said second component beingpositioned to receive radiation which has penetrated through a portionof said first component, and (c) means coupled to said first and secondcomponents to produce electrical signals representing radiation whenincident respectively on said first and second components.
 25. Thedetector of claim 24, wherein:said filter comprises material having anatomic number in the range of 24-58.
 26. The detector of claim 24,wherein:said first component comprises a phosphor layer comprising anelement having an atomic number lying in the range of 39-57.
 27. Thedetector of claim 24, wherein said second component comprises:a phosphorlayer comprising an element having an atomic number lying within therange of 56-83.
 28. The detector of claim 24, wherein one of said firstand second components comprises:a phosphor layer proximate and alignedwith a layer of light sensitive film.
 29. The detector of claim 24,wherein one of said first and second components comprises:a phosphorlayer proximate and aligned with a portion of photoconductive plate. 30.The detector of claim 24, wherein one of said first and secondcomponents comprises:a phosphor layer proximate and aligned with aportion of thermoluminescent plate.
 31. The detector of claim 24,further comprising:a filter of said penetrative radiation interposedbetween said first and second components.
 32. The detector of claim 31,wherein:said filter comprises material having an atomic number in therange of 24-58, and a thickness in the range of about 0.2 mm to 1.0 mm.33. the detector of claim 24, wherein said second material comprisesmaterial having a primary radiation absorber having a higher atomicnumber than that of said first material.
 34. The detector of claim 24,wherein:(a) said first material comprises one of yttrium oxysulfite,zinc cadmium sulfide, barium sulfate, barium cadmium sulfate, lanthiumoxysulide and barium fluorochloride, (b) said second material comprisesone of gadolinium oxysulfide, cadmium tungstate, calcium tungstate andbarium lead sulfate.
 35. The detector of claim 24, wherein:(a) saidfirst material comprises a first layer of phosphor material having acoating weight of about 20 to 100 mg/cm², and (b) said second materialcomprises a second phosphor layer having a coating weight of about 50mg/cm² to 1000 mg/cm².
 36. The detector of claim 24, wherein:(a) saidfirst component comprises a portion of a first scintillator material,and (b) said second component comprises a portion of a secondscintillator material.
 37. The detector of claim 24, furthercomprising:a portion of penetrative radiation filtering materialinterposed between said first and second components and being capable ofabsorbing substantially all radiation incident on said filter elementlying within said first energy range, while not absorbing substantiallyall such radiation of said second energy range.
 38. A method fordetecting area distribution of differing energy levels of penetrativeradation, said method comprising:(a) detecting preferentially lowerenergy radiation by passing it through a first detector elementincluding a scintillator and a plurality of segments; (b) detectinghigher energy radiation by transmitting radiation emergent from saidfirst detector incident onto a second detector element including ascintillator and a plurality of segments; (c) filtering penetrativeradiation emergent from said first detector before said second detectingstep, and (d) producing information in said first and second detectingsteps spatially locating radiation over an area with respect to at leastone of said detector elements.
 39. An energy discriminating radiationdetector comprising:(a) a first component comprising a first phosphormaterial including a primary radiation absorber having an atomic numberlying the range of 39-57; (b) a second component comprising a secondphosphor material aligned with said first phosphor material to receiveradiation when said radiation has penetrated through a portion of saidfirst component, said second phosphor material including a primaryradiation absorber having an atomic number lying within the range of56-83, and (c) means coupled to said first and second components forproducing electrical signals representing radiation when incident onsaid detector.
 40. An energy discriminating radiation detectorcomprising:(a) a first component comprising a first material of a firstkind which is preferentially responsive to penetrative radiation of afirst energy range; (b) a second component comprising a second materialdifferent in kind from said first material and of a kind which isresponsive to penetrative radiation of a second energy range extendinghigher than said first energy range, said second component being alignedwith said first component to receive radiation when said radiation haspenetrated through a portion of said first component, and (c) meanscoupled to said first and second components to produce electricalsignals representing penetrative radiation when incident on saiddetector.
 41. A radiation imaging system comprising:(a) a source forpropagating penetrative radiation along a path from a focal spot; (b) adetector assembly spaced from said source and interposed in said path,said detector assembly comprising:(i) a front array of individualdetector elements, each front array element including a penetrativeradiation sensitive receiving face having a discrete geometry, saidfront array element faces being located at substantially a distance F₁from said focal spot; (ii) a rear array of individual detector elements,each said rear array element including a penetrative radiation sensitivereceiving face having a discrete geometry, wherein each element of saidrear array is substantially aligned behind a corresponding element ofsaid front array, with respect to said focal spot, and in which eachrear array element has a receiving face which has a larger area then thereceiving face of its corresponding aligned front array element, saidrear array receiving faces being located at substantially a distance F₂from said focal spot, and (c) circuitry coupled to said detector arraysfor producing a representation of radiation when incident on saiddetector elements.
 42. The system of claim 41, wherein:(a) saidreceiving faces of said detector elements of said front and rear arrayshave similar geometry, and (b) a dimension D₁ of one of said front arrayelements is related to a corresponding dimension D₂ of one of said reararray elements by the following relation:

    D.sub.2 /D.sub.1 =F.sub.2 /F.sub.1.